Laser Surface Texturing of Polymers for Biomedical - Frontiers

REVIEW published: 27 February 2018 doi: 10.3389/fphy.2018.00016

Laser Surface Texturing of Polymers for Biomedical Applications
Antonio Riveiro 1*, Anthony L. B. Maçon 2, Jesus del Val 1,3, Rafael Comesaña 4 and Juan Pou 1,3
1 Applied Physics Department, University of Vigo, Vigo, Spain, 2 Department of Materials, Imperial College London, London, United Kingdom, 3 Department of Mechanical Engineering, Columbia University, New York, NY, United States, 4 Materials Engineering, Applied Mechanics and Construction Department, University of Vigo, Vigo, Spain

Edited by: Narayana Rao Desai, University of Hyderabad, India
Reviewed by: Krishna Chaitanya Vishnubhatla, Sri Sathya Sai Institute of Higher
Learning, India Khoi Tan Nguyen, Vietnam National University of HCMC,
Vietnam Sai Santosh Kumar Raavi, Indian Institute of Technology
Hyderabad, India
*Correspondence: Antonio Riveiro
[email protected]
Specialty section: This article was submitted to
Optics and Photonics, a section of the journal
Frontiers in Physics
Received: 13 November 2017 Accepted: 09 February 2018 Published: 27 February 2018
Citation: Riveiro A, Maçon ALB, del Val J, Comesaña R and Pou J (2018) Laser Surface Texturing of Polymers for
Biomedical Applications. Front. Phys. 6:16.
doi: 10.3389/fphy.2018.00016

Polymers are materials widely used in biomedical science because of their biocompatibility, and good mechanical properties (which, in some cases, are similar to those of human tissues); however, these materials are, in general, chemically and biologically inert. Surface characteristics, such as topography (at the macro-, micro, and nano-scale), surface chemistry, surface energy, charge, or wettability are interrelated properties, and they cooperatively influence the biological performance of materials when used for biomedical applications. They regulate the biological response at the implant/tissue interface (e.g., influencing the cell adhesion, cell orientation, cell motility, etc.). Several surface processing techniques have been explored to modulate these properties for biomedical applications. Despite their potentials, these methods have limitations that prevent their applicability. In this regard, laser-based methods, in particular laser surface texturing (LST), can be an interesting alternative. Different works have showed the potentiality of this technique to control the surface properties of biomedical polymers and enhance their biological performance; however, more research is needed to obtain the desired biological response. This work provides a general overview of the basics and applications of LST for the surface modification of polymers currently used in the clinical practice (e.g., PEEK, UHMWPE, PP, etc.). The modification of roughness, wettability, and their impact on the biological response is addressed to offer new insights on the surface modification of biomedical polymers.
Keywords: laser surface texturing, surface modification, wettability, surface roughness, implants, cell response
INTRODUCTION
Polymers are organic materials, formed by linking a large number of repeating units called monomers. These materials are widely used in biomedical applications, e.g., in joint replacement components. Typical polymers used in clinical applications include polyetheretherketone (PEEK), ultra-high-molecular-weight polyethylene (UHMWPE), polypropylene (PP), acrylic bone cements (PMMA), or nylon among others. They exhibit excellent mechanical properties for applications such as knee and hip implants, sutures, orthopedic fixation implants (pins, screws, rods, clips, etc.), dental implants, or stents among others. In addition, they have a reduced density as compared to other biomaterials (such as metals or ceramics), and do not interfere and degrade the biological tissue in contact.
Although these materials are biocompatible, and in some cases, have similar mechanical properties to human tissues, they are, in general, chemically and biologically inert. They show

Frontiers in Physics | www.frontiersin.org

1

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

a minimum interrelation with the surrounding tissues or cells. In consequence, the body usually responds by forming a nonadherent fibrous tissue (through a process called fibrosis) around the surface of the material, and progressively, they are completely encapsulated by such layer. Several works have reported the severe reduction in the mechanical strength of a fibrous encapsulated implant compared with a properly osseointegrated implant [1]. On the other hand, these problems can be worsened by bacterial, viral, or fungal infections occurring around the implant. In this case, the removal of the implant (so-called revision surgery) may be the best alternative, because the fibrous tissue can be impermeable to the medications [2]. One strategy to solve these problems is the application of a bioactive surface coating (e.g., bioactive glass, hydroxyapatite, titanium dioxide, etc.), or to mix these polymers with bioactive materials (e.g., bioactive glass, hydroxyapatite, β-TCP, calcium silicate, etc.) [3– 6]. The main problem of coatings is to guarantee a proper adhesion to the polymeric surface. Furthermore, the mixture of polymers with bioactive materials can drastically reduce their mechanical properties. One alternative consists in the application of a surface treatment to enhance the biological properties of the polymers without compromising their mechanical properties. In this way, the fibrosis of implants can be effectively reduced, promoting the tissue integration.
ROLE OF MATERIAL SURFACE
TOPOGRAPHY AND WETTABILITY ON THE
BIOCOMPATIBILITY
Biocompatibility is intimately related to the response of cells in contact with the surface of a given material, and in particular with their adhesion [7]. The response of tissues to an implant mainly depends on the physico-chemical properties of its surface. Surface properties such as topography (or texture), surface chemistry, surface energy, or wettability determine the interaction of implants with the biological environment.
The main aim of surface texturing techniques in biomedical applications is the enhancement of the cellular activity in the surface of the implant. In bone remodeling, textured surfaces show a higher surface area for integrating the implant with bone, via osseointegration process. Furthermore, textured surfaces also allow ingrowth of the tissues, and promotes the mechanical stability of implants [8]. Different length scales can be distinguished in an implant surface. Macro-topographies, with surface roughness ranging from millimeters up to microns, can contribute to improve the fixation and long-term mechanical stability of the implant device [9]. Micro-sized topographies, in the range of the microns, affect cell adhesion and proliferation (see Figure 1); they have a well-established influence on the improvement of the osseointegration of implants [10]. This kind of features play a key role in the adsorption of proteins. They also affect the cell proliferation, and enhance cell adhesion, and it has been determined their influence on the gene expression for different cell types; therefore, they could be potentially used as a signaling modality for directing differentiation [11].

Another physicochemical property of the implant surface is the interfacial free energy (or in short, surface energy) of the material. This parameter is closely related to the wettability (the preference of a solid to be in contact with one fluid rather than another) of the material. This property is typically evaluated with the water contact angle (WCA), i.e., the angle formed by the interface liquid-vapor with a solid surface (as depicted in Figure 2). When a material has a high affinity for water (hydrophilic), i.e., high surface energy, the water spreads on the material and the contact angle is low. In the opposite case (hydrophobic), i.e., low surface energy, water does not spread and forms, at equilibrium, a spherical cap resting on the surface of the material with a large contact angle (see Figure 2). It is shown that, more hydrophilic substrates (i.e., with high surface energy, low contact angles) promotes considerably the adhesion and spreading of cells as compared to in hydrophobic materials (i.e., with low surface energy, high contact angles). Air bubbles trapped on the surface of hydrophobic surfaces prevent the protein adsorption to surfaces, and subsequent interaction with cell receptors [10]. This avoids the normal cell adhesion on hydrophobic surfaces (see Figure 1). In this sense, Schakenraad et al. [12] demonstrated that the cell spreading is higher on the surface of hydrophilic materials than on hydrophobic (in absence of preadsorbed serum proteins). However, cell adhesion can decrease if the material becomes excessively hydrophilic. This suggest the existence of an optimum range of surface energies [13].
Finally, we should point out that surface topography and wettability of any material are not unrelated properties. Surface roughness can directly affect the wettability of materials and arises as an effective method to control this property. Then, both parameters modulate the biological response.
One of the first attempts to determine the influence of the roughness of any surface on its wettability characteristics was due to Wenzel [14]. He demonstrated the promotion of the intrinsic wettability of a material with the surface roughness. Therefore, if the surface is chemically hydrophobic, it will become even more hydrophobic when surface roughness is added, and vice versa if it is chemically hydrophilic. The results given by Wenzel are based on the assumption that the liquid penetrates into the roughness grooves (see Figure 3). In cases where the liquid does not penetrate into the grooves, the Wenzel model does not longer applies. Cassie and Baxter addressed this problem, and assumed that the liquid does not penetrate into the grooves, as depicted in Figure 3 [15]. The main result of this model is that the presence of air pockets tends to increase the hydrophobicity of the material, independently of its intrinsic wettability. It is generally observed that very rough surfaces are more likely to follow the CassieBaxter regime, and low rough surfaces will follow Wenzel’s model [16]. Therefore, it can be deduced from this analysis that we can tailor the wettability of the material, and then its biological performance just by changing its surface roughness.
Another standard method to modify the wettability and biological performance of a material is by varying its surface chemistry. Materials with non-polar surfaces, such as polymers, have low surface energy. The modification of the surface chemistry to induce the presence of polar or charged functional

Frontiers in Physics | www.frontiersin.org

2

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

FIGURE 1 | Influence of the wettability and surface roughness on the cell adhesion.

FIGURE 2 | Water contact angle for hydrophilic (i.e., with high surface energy), and hydrophobic materials (i.e., with low surface energy).

FIGURE 3 | Schematic of the influence of the roughness on the wettability. The Cassie–Baxter model proposes that water droplets sitting on rough surfaces form a solid–air–water interface. Air pockets are trapped beneath the droplet. In the Wenzel model, no air pockets form, and the surface is completely wetted by the droplet.

groups is a technique used to increase the wettability of polymers [17].
BASICS OF LASER SURFACE TEXTURING
Introduction
Surface modification techniques of polymeric biomaterials for medical implants are performed, in general, in two different ways to promote their biological characteristics. These techniques rely on: (1) deformation, removal or controlled addition of material to the surface to increase the roughness, or (2) by modifying its surface chemistry [18, 19]. Techniques such as photolithography, focused ion beam micromachining, direct writing techniques, or transfer printing, among others, are able to modify polymeric surfaces at the micro- and nanoscale. However, the potentiality of these methods does not cover the

full spectrum of requirements needed for the direct treatment of current polymeric biomaterials used in implants [20]. Most of them involve the utilization of toxic chemicals, require multiple steps, sterilization is not guarantee along the treatment (being needed a post-sterilization stage), and production of hierarchical structures is not always possible (see Table 1). An alternative technique to these methods, called laser surface texturing (LST; laser texturing, laser structuring, or laser patterning), is based on the direct treatment of polymeric biomaterials with a laser beam. This technique offers a great number of advantages; in particular, the most important is the possible modification of surface roughness and chemistry in one step avoiding the utilization of toxic substances. Laser surface texturing can modify polymeric surfaces at a macro-, micro-, and nano-size scale with a high spatial and temporal resolution [21]. Given the non-contact nature of the process, the contamination of the

Frontiers in Physics | www.frontiersin.org

3

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

TABLE 1 | Characteristics of some processing techniques used to produce surface modifications on polymeric biomaterials.

Technique

Processing Processing Chemical Treated Cost

rate

steps

products area

Photolitography Electron beam litography Ion beam litography Atomic force microscopy Soft lithography Chemical vapor deposition Laser texturing

Fast Slow
Slow
Slow
Fast Fast
Fast

Several

Yes

One

No

One

No

One

No

Several

Yes

One

Yes

One

No

Large High Small High
Small High
Small High
Large Low Large High
Large Medium

workpiece is easily avoided; this is a very important advantage for biomedical applications as the sterilization of the implants can be guaranteed. Another advantage is the high processing speed, the easy automation, and the possibility to treat large areas. The simultaneous modification of roughness and chemistry also leads to the simultaneous change of the wettability (or surface energy) of polymers.
The potentiality of this approach to enhance the biological response of biomaterials was largely studied in metals (in particular in titanium alloys) [22–24], and more recently in polymeric biomaterials. Most of the research work done on polymeric biomaterials in this regard only evaluate the biocompatibililty in terms of the change in roughness and wettability with the laser treatment. More recently, some research works (using in vivo and in vitro tests) have been performed to elucidate this enhancement in the biological response of the laser-treated polymeric biomaterials; however, more studies are still needed to transfer this technique into the current clinical practice. With this work, we intend to complement reviews on laser patterning, mainly concentrated on laser ablation mechanisms (e.g., [21, 25–27]), and to provide the reader with a critical understanding on the next steps to advance in this field.
Process Fundamentals
LST is one of the simplest techniques to modify both surface topography and chemistry [28]. In this case, a focused laser beam is directed onto the surface of some material; then, the laser radiation is absorbed by the topmost layer (Figure 4). The optical energy provided by the laser beam induces the heating of the material, reaching the melting, or even the vaporization temperatures. This way a selective material removal is achieved, and the surface topography is modified. On the other hand, if the photons of the laser beam are sufficiently energetic, e.g., using UV-lasers, they are able to break chemical bonds, and then modifying the surface chemistry of the material. Therefore, thermal and/or photochemical processes can modify the surface of polymers:
• Thermal processes: the temperature of the material is increased by the thermalization of the optical energy in the

surface of the material. This phenomenon leads to diverse induced phenomena, including melting or vaporization [29]. In general, these phenomena induce the modification of the surface roughness. • Photochemical processes: the energy of photons emitted by the laser source is so high that directly breaks the molecules of the treated surface. This mechanism is the main responsible for the chemical modification of surfaces. In this case, due to the necessity of high energy photons, ultraviolet (UV) lasers are the most commonly employed ones [30]. • Photophysical processes: in this case, thermal and photochemical process jointly influence on the process [31]. In this case, both surface roughness, and chemistry can be simultaneously modified.
LST can be performed by the creation of regular or irregular patterns of bumps, dimples, and (linear or non-linear) grooves as depicted in Figure 4 [28]. The resulting surface topography and chemistry depends on the preponderance of thermal, or nonthermal processes. If the surface of the material is melted, bumps or dimples can be formed due to the formation of projections, depressions or due to the foaming of the material. The increased absorption of laser radiation can produce the vaporization of the material. In this case, removal of material is mainly produced by the vaporization or thermal decomposition, but some melting or thermal degradation can also occur. Non-thermal processing associated to the utilization of ultrafast lasers can also produce dimples and grooves, but avoiding undesirable thermal effects (e.g., generation of a heat affected zone). In this case, the direct breaking of molecular or atomic bonds takes place, rather than simply heating. Therefore, this is a clean process, leaving no recast material and eliminating the need of post-processing steps [31].
Patterns in biomedical polymers can be produced in the UV–IR spectral range, using continuous-wave (CW) or pulsed laser radiation [32, 33]. Influence of the processing parameters, their effects, and theoretical modeling is complex. It depends on multiple parameters associated to the nature of the polymer and the specific working approach used to produce the surface topography (e.g., photo-thermal or photo-chemical ablation of the material, laser swelling or bumping, laser grooving, etc.). These analyses are beyond the scope of this work. We refer the interested reader to the specialized works for further information (e.g., [21, 34–41]).
Absorption characteristics of polymers depends on their structure, but this behavior is influenced by the presence of fillers or additives. UV radiation is preferred for LST as seen in Table 2, but other laser wavelengths are also used. IR radiation tends to produce the thermal ablation or melting of polymers, while the treatment with UV radiation is able to ionize and decompose polymers without substantial melting. This laser radiation can also modify the surface chemistry of polymers. In this case, the polar component of the surface energy can be considerably increased, and the wettability can be promoted (see e.g., [42]).
Compared to pulsed laser radiation, CW operation mode produces patterns with low quality, and high thermal affectation

Frontiers in Physics | www.frontiersin.org

4

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

FIGURE 4 | Schematic of the principle of operation of laser surface texturing (LST).

around the laser-treated area. Then, this processing mode is scarcely used (see Table 2). On the contrary, the reduction in the pulse length produces high precision patterns. In the case of processing with ultrashort (picosecond or femtosecond) laser pulses, the heat diffusion into the polymer is avoided, and the thermal damage close to the radiated area is negligible [76]. Therefore, debris formation can be avoided. This is interesting as any potential source of debris in implants must be avoided. Released implant debris induces inflammation and osteolysis around the implanted area, and compromises the performance of implants [77]. Another relevant feature about using ultrashort pulses is the high intensities produced in the irradiated area. These are able to produce nonlinear optical phenomena, such as multiphoton absorption. Then, molecular bond breaking or even ionization can be reached even though the wavelength of the laser radiation is theoretically long (i.e., low photon energies) [78].
Other laser and non-laser related processing parameters are listed in Table 2. These also depend on the nature of the polymer. In general, focused laser beams onto the surface of the samples are used, with spot diameters ranging from several to hundred (even thousands) of microns. LST experiments are predominately performed in air atmosphere (see Table 2); however, the influence of the processing atmosphere should not be neglected as this affects the surface chemistry and the wettability of the laser-treated surfaces. Pfleging et al. [71] observed a marked increase in the wettability of PMMA using O2 (instead of He) as processing gas during laser processing of PMMA. XPS measurements showed the oxidation of the PMMA surfaces during the processing with O2. However, the influence of this processing parameter on the biological performance of laser textured samples has not yet been evaluated, and this evaluation should be addressed in future studies.

Components of a Laser Texturing System
There are two main methods for laser texturing: (1) using a stationary laser beam, or (2) providing a relative movement between the laser beam and the surface of the workpiece.
The first approach requires the utilization of a mask with the desired pattern. Then, only the portion of light that passes through the mask is imaged by a lens to produce a pattern on the surface of the workpiece (see Figure 5). In this case, a short laser pulse (required to obtain high peak powers) is used to produce the pattern; then, pulsed lasers (such as excimer, femtosecond, or TEA CO2 lasers) are utilized. This approach is not commonly used (see Table 2) as it is less flexible. Production of a mask is a time-consuming process, and the modification of the textured pattern necessarily requires the production of new masks. Furthermore, some small structures can be formed around the textured pattern due to the laser beam diffraction at the mask geometry [71]. Other approach also using a stationary laser beam is the laser interference patterning, or direct laser interference patterning (DLIP) [70]. This technique involves the interference of two or more laser beams to obtain a periodic variation of light intensity on the irradiated area. This approach can produce very fine details, and even the production of hierarchical structures with two different length scales. This posibility can be very attractive to promote osteoconductive and osteoinductive properties because the cell response (morphology, adhesion, proliferation, etc.) depends on the micro- and nanotopography [79]. However, this technique is less flexible than using a mask: (1) the modification of the pattern is not simple (the production of a defined interference pattern requires a complex and time consuming procedure to superimpose the laser beams with high accuracy), and (2) the system is more expensive [80].
In the second case, the relative motion of the laser beam with regard to the workpiece produces the pattern. This can be done moving the beam or the workpiece. Two common ways

Frontiers in Physics | www.frontiersin.org

5

February 2018 | Volume 6 | Article 16

Riveiro et al.

Frontiers in Physics | www.frontiersin.org

TABLE 2 | Summary of processing parameters and research aims of selected works found in the literature on LST of clinically relevant polymers.

Material Form/

Laser

λ (nm) Power Spot Fluence Focal CW/P

Freq. (Hz)

Pulse Pulse Texturing approach & Scanning Atmosphere Application

Ref.

thickness

(W) diameter (J/cm2) length

duration number morphology

speed

(µm)

(mm)

(mm/s)

6

PEEK PEEK
PEEK PEEK

Film 250 µm ArF excimer

Plate 8 mm Nd:YVO4

Nd:YVO4

Nd:YVO4

–

Nd:YAG

–

ArF excimer

193 –

–

5-80x10−3

1,064 5–39 100–1,000

–

532 5–39 100–1,000

–

355 5–39 100–1,000

–

355 –

450

0.75

193 –

–

3.5

–P
211 P
365 P 235 P
–P
–P

10
15,200–50,000
15,200–50,000 15,200–50,000
–
–

20 ns
∼ns
∼ns ∼ns 38 ns
20–25 ns

1–2,000
–
– – 4–12
10

Stationary beam/Conical structures
Galvanometric scanning system/Parallel grooves
Idem
Idem
Interference patterning/Line structures
Stationary beam + Mask/Parallel grooves

–

Ar, Air, O2

200–5,000 Air

200–5,000 Air

200–5,000 Air

–

–

–

–

PEEK PEEK PEEK

Membranes CO2 240 µm

10,600 1

Plate 3.92 mm

Q-Switched Nd:YAG

1,064 –

Foil 50 µm KrF Excimer 248 –

–

–

600–2,500 0.01–0.1

–

0.004–0.03

100 Quasi-CW –P –P

5,000 10 10

∼ms 10 ns

– Galvanometric scanning system/ Parallel grooves
– XY Table/Dimples

20–40 ns 6,000 XY Table/Ripple & dot patterns

–

Air

12 Air

–

–

PEEK

Sheet 2 mm Ti:sapphire

800 0.02

–

–

–P

1,000

–

– Parallel grooves

1 Air

PEEK+ Nano-SiO2 UHMWPE
UHMWPE

Sheet 2 mm Ti:sapphire

Sheet 1 mm Sheet 1 mm

Q-Switched Nd:YAG Q-Switched Nd:YAG KrF excimer XeCl excimer Nd:YAG

800 0.02
1,064 –
532 –
248 – 308 – 532 –

UHMWPE –

CO2

10,600 –

–

KrF excimer 248 –

UHMWPE UHMWPE

Disk 3 mm Plate 8 mm

Yb:KYW femtosecond
Nd:YVO4

1,027 – 1,064 0.8–6

PP

–

Nd:YVO4 Nd:YVO4 ArF excimer

532 1.5-6.5 355 0.3–1.4 193 –

–
–
–
– – 6,000
–
–
8
70
60 25 –

–
0.6
0.6
0.6 0.6 0.5
<2.5
0.3
2–14
<8.3
<11.4 <11.4 0.0125

–P
–P
–P
–P –P No lens P
–P
–P
–P
211 P
365 P 235 P
–P

1,000

–

– Parallel grooves

1 Air

0.5

9 ns 10–50 Stationary beam/Rough

–

Air

surface

0.5

9 ns 10–50 Idem

–

Air

0.5 0.5 10
–
–
1,000

23 ns 10–50 Idem

–

Air

20 ns 10–50 Idem

–

Air

3 ns

0 Stationary beam/ Rough

–

Air

surface

30 ns

20–50 Stationary beam/Rough surface

–

Air

23 ns

20–50 Stationary beam/Rough surface

–

Air

450 fs

– XYZ Table/ Rough surface 1–2 Air

15,200–50,000
15,200–50,000 15,200–50,000
–

∼ns

– Galvanometric scanning 200–5,000 Air

system/Parallel grooves

∼ns

– Idem

200–5,000 Air

∼ns

– Idem

200–5,000 Air

– 0–10,000 Stationary beam/Surface

–

Tap water

PP

–

Ti: sapphire

800 –

60

0.4–3.2

–P

5,000

130 fs 1-50 XY Table /Craters

–

Air

PP

–

Nd:YAG

1,064 –

7,000

5

–P

10

2-8 ns

––

–

Air,O2 ,N2

Nd:YAG

532 –

–

2.5

–P

Nd:YAG

266 –

–

0.45

–P

PP

Sheet 5 mm KrF excimer 248 –

–

5

–P

1–2 ns

––

–

Air,O2 ,N2

3–4 ns

––

–

Air,O2 ,N2

1

–

– Stationary beam/Rough

–

Air

surface

Influence on adhesive

[43]

bonding properties

Biomedical (Increment of [44] roughness and wettability)

Biomedical (influence on cell [45] growth)

Biomedical (Influence on [46] pre-osteoblastic cell response)

Biomedical (Influence on [47] osteoblast response)

Influence on adhesive

[48]

bonding properties

Biomedical (Influence on [49] wettability, roughness, chemistry and cell adhesion)

Biomedical (Influence on [50] osteogenic response)

[50]

Influence on wettability

[51]

Influence on roughness

[52]

Influence on roughness, [53] wettability and chemistry

Influence on surface

[54]

topography and chemistry

Biomedical (Influence on [55] roughness and wettability)

Influence on bonding

[56]

characteristics

Influence on surface

[57]

properties

Influence on bonding

[58]

characteristics

Biomedical (Influence on [59] roughness, wettability, cell proliferation)

LST of Biomedical Polymers: Review

February 2018 | Volume 6 | Article 16

(Continued)

February 2018 | Volume 6 | Article 16

7

Frontiers in Physics | www.frontiersin.org

TABLE 2 | Continued

Material

Form/

Laser

thickness

λ (nm) Power (W)

Spot diameter
(µm)

Fluence (J/cm2 )

Focal CW/P length (mm)

PE Low-density PE PE

Film 300 µm Film 50 µm
Sheet 2 mm

Nd:YAG TEA CO2
Ti:sapphire

532 – 9,580 –
790 –

1,100 –
1.5

1.4–6.3 1.2
<0.11

300 P –P
150 P

Ti:sapphire (second harmonic)

395 –

1.5

<0.08

150 P

395 + – 790

1.5

<0.08

150 P

ArF excimer 193 –

–

<0.08

–P

Molecular fluorine (F2)

157 –

–

<0.02

–P

PE

Sheet 2 mm Nd:YAG

1,064 –

–

–

–P

Bisphenol-A Thin film KrF excimer 248 –

PC

3 µm

PC

Film

ArF excimer 193 –

–

<0.35

–P

–

0.01–0.02

–P

248 –

–

0.01-0.03

–P

PC

Film 1 mm Nd:YAG

1,064 –

0.8-3 0.42-19.9

–P

Nylon 6,6 Nylon 6,6

Sheet 5 mm CO2 Film 1 mm CO2

355 –

3

10,600 7

95

10,600 7–10

–

0.42

–P

–

250 CW

–

––

PI PI
PMMA

Film 3 µm Nd:YAG

355 –

Film >2 mm Nd:YAG

355 –

Nd:YAG

266 –

–

ArF excimer 193 –

–

0.1-1

–P

–

0.1-3

––

–

0.1–3

––

–

0.02–0.09

–P

PMMA

Sheet 3 mm Ti:sapphire

PMMA

Film 1 mm Ti:sapphire

PMMA PMMA

Film 1 mm Film 1 mm

Ti:sapphire Ti:sapphire

CW, Continuous wave; P, Pulsed.

800 –
800 –
775 – 800 –

20.4
1.5–2.4
29 1.5-2.4

0.85–50
–
0.1-50 –

100 P
–P
100 P –P

Freq. (Hz)

Pulse Pulse Texturing approach & Scanning Atmosphere Application

Ref.

duration number morphology

speed

(mm/s)

1

10 ns 1–300 Stationary beam /Dimple

–

–

0.5

–

1–3 Stationary beam/Rough

–

Air

surface

–

130 fs

– Stationary beam +

Mask/Dimple

–

Air

–

130 fs

– Idem

–

Air

–

[60]

Influence on surface

[61]

properties

Influence on surface

[62]

properties

– 10 10 500 1 10 10 1 1 – –
10 10 10 –
1,000 1,000 1,000 1,000

130 fs

– Idem

23 ns 20 ns

– Idem – Idem

–

––

30 ns 20 ns

5 Stationary beam/Rough surface
1–1,000 X–Y Table

30 ns 10 ns

1–1,000 Idem 1-5 Rotating table/Craters

10 ns –
–

1-5 Idem
– Galvanometric scanning system/ Parallel grooves
– Galvanometric scanning system/Parallel grooves

10 ns 10 ns 10 ns

– Interference patterning/Line structures
– Interference patterning/ Dot & line structures
– Idem

20 ns 1-1,200 Grooves

150 fs
100 fs
∼fs 100 fs

20 XY Table/Parallel grooves
- XYZ Table/Craters
– Parallel grooves – XYZ Table/Craters

–

Air

–

Air

–

Air

1–

Biomedical (Influence on [63] biological response)

–

Air

Biomedical (Influence on [64] roughness)

–

Air, Ar

Biomedical (Influence on [65] chemistry, wettability)

–

Air, Ar

-

Air

Biomedical (Influence on [66] fibroblast response)

–

Air

600 Air

Biomedical (Influence on [67] fibroblast response)

600 Air

Biomedical (Influence on [68] roughness, wettability, cell response)

–

Air

Biomedical (Influence on cell [69] response)

–

Air

Influence on surface

[70]

topography

–

Air

Influence on surface topography

–

Air, O2, He Biomedical (Influence on [71]

wettability, chemistry, cell

adhesion)

1 Air

Influence on surface

[72]

properties

–

Air

Influence on surface

[73]

topography

2

Microfluidics

[74]

1 Air

Microfluidics

[75]

LST of Biomedical Polymers: Review

Riveiro et al.

Riveiro et al.

LST of Biomedical Polymers: Review

of doing this are: the utilization of a Cartesian system to move the laser beam (see Figure 6A), or steering the beam over the workpiece via two galvanometer mirrors (see Figure 6B). The last approach is more commonly selected due to the higher versatility, and throughput. In both systems, CW or pulsed lasers can be used.
FIGURE 5 | Laser surface texturing using a stationary beam and a mask.

As seen in Table 2, different laser sources are used for LST. UV laser radiation is generated by excimer (e.g., ArF or KrF), or solidstate lasers (e.g., frequency tripled); however, the better beam quality of solid state lasers, and the non-utilization of toxic gases (in contrast with excimer lasers) make solid-state lasers more attractive for LST. Nd:YAG or Nd:YVO4 lasers emitting ns laser pulsed in the NIR range are also used. Ti:sapphire lasers, emitting 800 nm laser radiation, were used to produce fine patterns. Far-IR laser radiation from CO2 lasers is barely used.
LASER SURFACE TEXTURING OF
BIOMEDICAL POLYMERS
Multiple polymeric biomaterials have been studied for tissue engineering [81–83]; however, only few of them are really used in the current clinical practice. Therefore, we will review the main research works on LST of those currently used for medical applications, especially on those used in bone tissue engineering [84].
Poly(etheretherketone) (PEEK)
Polyetheretherketone (PEEK) is an engineering thermoplastic that exhibits a rigid semicrystalline structure. It shows excellent mechanical properties (in some cases, even similar to cortical bone), and high chemical resistant [85]. In addition, PEEK exhibits sterilization capacity [86]. These properties turn it into an ideal material to be used in biomedical applications. This polymer is mainly used as structural material in orthopedic applications (e.g., in joint replacement, cage implants, bone screws, and pins, etc.); however, it is biologically inert due to its large chemical stability, and low wettability (see Table 3). This leads to poor bone-implant interactions.
First studies on LST of PEEK were performed by Laurens et al. using excimer lasers [43]. It was demonstrated the ability of ArF laser (λ = 193 nm and pulse duration = 20 ns) to modify PEEK surfaces below the ablation threshold. The chemical modification after laser treatment depended on the assist gas used in the process (i.e., a neutral or an oxidizing atmosphere). Under neutral conditions, loss of polymer aromaticity and scission of carbonyl groups took place. In the presence of

FIGURE 6 | Laser surface texturing by imposing a relative movement between the laser beam and the workpiece using (A) a cartesian, or (B) a scanning system.

Frontiers in Physics | www.frontiersin.org

8

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

TABLE 3 | Surface roughness and contact angle of pristine polymeric biomaterials, and after laser surface texturing experiments.

Material

Laser/wavelength (nm)

Ra (µm)

θ(◦ )

Base material

Ra (µm)

θ(◦ )

Laser treated surface

PEEK PEEK UHMWPE UHMWPE UHMWPE Nylon 6,6 PP PP

KrF excimer/248 Ti:Sapphire/800 KrF excimer/248 Nd:YVO4 /1,064 Nd:YAG/1,064 CO2 /10,600 KrF excimer/248 Nd:YVO4 /1,064

1.4·10−3 0.3193 0.075 2.3 ± 0.4
– 0.29 0.0163 0.057

69.4 79.3 94 82 ± 5 94 66.49 ± 0.32 99 86.1 ± 2.0

17·10−3 0.6245
0.6 2.3 ± 0.7
– 0.83 0.02023 4.04

50 54.8 86 75.2 ± 10 90 57.5 ± 2.50 76 89.9 ± 13

References
[49] [50] [53] [55] [51] [68] [59] [87]

environmental oxygen, the high energetic incoming photons increased the C-O/C, and carboxylic functions. Furthermore, the polar component of the work of adhesion increased after laser treatment. Similar results were also observed by Michaljanicˇová et al. [49] during the treatment of PEEK surfaces with UV radiation (λ = 248 nm). In this case, an increase of the wettability is observed (see Table 3). This was attributed to the increase in roughness, and oxygen containing groups formed on the treated surfaces.
More recently, other laser wavelengths were studied with the aim to evaluate their effect on the surface modification of PEEK [44, 47, 48]. So far, surface functionalization of PEEK by means of LST has been successfully achieved using laser wavelengths ranging from UV (355 nm) to middle infrared (MIR) (10.6 µm).
Riveiro et al. compared the effect of the three ns laser irradiation wavelengths (355, 532, and 1,064 nm) on the roughness, and contact angle of PEEK substrates, properties directly related to the cell viability on implants [44]. In order to enhance its adhesion properties, an increase on the wettability was sought. PEEK was observed to respond very differently as a function of the laser radiation (see Figure 7). The 1,064 m laser radiation burned the surface, while the 532 nm laser radiation was able to ablate the material. Using this laser wavelength, grooves with a mean width of 100 µm were machined. The 355 nm laser radiation only produced a slight surface melting; however, this laser radiation was identified as the most suitable for biomedical purposes because induced the formation of some polar groups [carboxyl (O-C=O) and peroxide (O-O)] on the surface of the PEEK samples. They produced a marked reduction in the WCA of PEEK after treatment. This fact can be potentially beneficial to promote cell adhesion onto laser treated PEEK. A similar effect was observed by exposing PEEK to Q-switched Nd:YAG laser radiation (λ = 1,064 nm) [48]. In this case, the surface energy was increased (from 44.9 to 78.5 mJ/m2), along with the wettability, after the laser treatment. Chemical analyses suggested that an increase in hydroxyl and carboxylic groups, along with a decrease on the original carbonyl groups, took place after laser treatment. It was showed the formation of functional polar groups on lasertreated PEEK surfaces. These findings confirmed the NIR lasers as a viable option to promote the functionalization of PEEK surfaces by LST.

These studies only assessed the biocompatibility of laser treated surfaces in terms of the wettability; however, this does not guarantee the biocompatibility of the surfaces, and biological tests (in vitro and also in vivo tests) are required.
The enhancement of biocompatibility of the laser textured PEEK surfaces attributed to the formation of polar groups was confirmed by Zheng et al. [47]. In vitro tests, along with topographical, and chemical analyses on PEEK surfaces modified by a combination of CO2 laser (λ = 10,600 nm) and plasma treatments, were conducted. MC3T3-E1 pre-osteoblast were the cells employed in this study. Cell adhesion and proliferation were increased after laser treatment, in conjunction with an increase in the formation of carboxylic groups on the surface. Therefore, it was established the direct relationship between the roughness and the formation of polar groups with the increased biocompatibility of laser-treated PEEK surfaces [47]. Guo et al. [50] focused their attention to the influence of the roughness on the biological activity and osteogenic efficiency of laser-treated surfaces. Femtosecond laser irradiation was used to modify the surface of PEEK implants (with and without the reinforcement of nano-SiO2 particles). In vivo animal tests were performed on rabbits, and demonstrated a superior bonding strength of the bone/implant interface for treated implants. A similar result was found by Briski et al. [88] during implantation of laser textured PEEK cages for fusion in a sheep model. Enhanced fusion and higher deposition of mineralized matrix was observed after 6 months of implantation.
Cordero et al. [46] studied the influence of micro-grooves patterns on the cell alignment of MC3T3-E1 pre-osteoblasts. Patterns were produced with UV light (from an ArF laser) using a mask projection unit. The increased cell growth along preferential directions is proposed to act as a bridge for the bone regeneration in disrupted areas. In vitro results, showed a better cell alignment for patterns formed with lines separated distances below 50 µm. Bremus-Koebberling et al. [45] followed a similar approach, but using nano-grooves. They evaluated the influence of linear-like nanopatterns produced by laser interference patterning (λ = 355 nm, pulse duration = 38 ns) on the cell alignment using B35 neuronal cells. In this case, the main aim was their utilization for neuronal repair applications. It was demonstrated that not only the pattern periodicity (i.e.,

Frontiers in Physics | www.frontiersin.org

9

February 2018 | Volume 6 | Article 16

Riveiro et al.

LST of Biomedical Polymers: Review

FIGURE 7 | SEM images of the laser treated PEEK surfaces under 1,064, 532, 355 nm laser wavelengths, and for the base material. Insets: optical images of water drops deposited onto untreated and laser treated PEEK surfaces (untreated PEEK: contact angle 92.4◦, 1,064 nm laser radiation: contact angle 59.5◦, 532 nm laser radiation: contact angle 131.0◦, 355 nm laser radiation: contact angle 52.5◦). Adapted from [44] with permission from Elsevier.

the width of the nano-grooves), but also the groove depth (or more precisely, the aspect ratio) influences the cell alignment. These results demonstrate that nano-topography is relevant for the control of the cellular response. They also open the door for the application of laser-textured patterns not only for bone tissue engineering, but also for other biomedical applications.
Ultra-High-Molecular-Weight Polyethylene
(UHMWPE)
Ultra-high-molecular-weight polyethylene (UHMWPE) is a thermoplastic synthetized from monomers of ethylene (-(CH2)n-), used as a bearing material in joint replacement implants for decades. It is commonly used in biomedical applications such as in patella or hip prostheses [51, 55]. This is principally due to its high wear and impact resistance, chemical stability, and non-toxicity in contact with corporal fluids. However, the large chemical stability, and low wettability (see Table 3) make this material also bioinert.
Laser surface modification on UHMWPE has been successfully demonstrated by means of fs, ps, and ns laser sources. Laser treatments produces the increment of the roughness and wettability (see Table 3).
Pulsed Yb:KYW fs laser (λ = 1,027 nm, and pulse duration = 450 fs) was used to create size-controlled craters on UHMWPE surfaces. The maximum ablation efficiency was obtained using laser pulse energies above 6 µJ. The surface chemistry of UHMWPE was found to remain virtually unaltered after laser interaction [54]. However, by subjecting this thermoplastic to ps laser surface radiation (from the iodine PALS -Prague Asterix Laser System- laser), different results were revealed [89]. In this case, 438 nm laser pulses, with a high energy (up to 240 J) and a short pulse length (400 ps) were used. The chemical analyses of the irradiated surface confirmed substantial changes in the ablated crater composition when compared to the pristine UHMWPE substrate. The results from the mass quadrupole

spectrometer (MQS) showed that C-H, and C-C chemical bonds were broken given the high deposited energy during the ablation process. Also, Raman spectroscopy and infrared absorption analyses demonstrated the enrichment in carbon content of the laser affected areas.
Lorusso et al. [51] evaluated the influence of the laser treatment on the wettability using IR (λ = 1,064 nm), visible (λ = 532 nm), and UV laser radiation (λ = 308 and 248 nm). They found a decrease of the WCA with the number of pulses (irrespectively of the laser wavelength). This reduction is more marked for surfaces treated with UV laser radiation. These results are explained in terms of the interaction mechanisms with the substrate. IR and visible laser pulses induces photothermal effects, while UV radiation is able to produce photochemical effects. It is postulated that they modify the surface chemistry. Velardi et al. [53] also found using 248 nm UV laser radiation (and similar processing parameters) that this laser wavelength produces the formation of oxidized groups (such as hydroxyl and carbonyl). A similar decrease in the WCA, and the increment of the average roughness with the number of pulses was also observed. Then, the increment of the wettability is mainly ascribed to the modification of the surface chemistry.
Riveiro et al. studied the modification of UHMWPE with several laser wavelengths (1,064, 532, and 355 nm) [55]. In this work, a carbon coating was applied on the surface of the UHMWPE samples to reduce the large transparency exhibited by this material to these laser wavelengths. Raman spectroscopy revealed that UHMWPE undergoes minor chemical modifications after the laser treatments. The visible and UV laser wavelengths were found the most suitable to tailor the topography, and wettability. In both cases, the average roughness was closer to Ra = 1 µm, and the WCA was reduced compared to the pristine surface (Figure 8). Nevertheless, due to the melting of the surface after the laser treatment, some carbon particles were trapped during the process (Figure 8);

Frontiers in Physics | www.frontiersin.org

10

February 2018 | Volume 6 | Article 16