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A Comparative Study of 316L Stainless Steel and a 

Titanium Alloy in an Aggressive Biological Medium  
 

Djelloul Aroussi 

Laboratory of Applied Biomechanics and 
Biomaterials (LABAB) 

ENPO-MA, Oran, Algeria 

djelloul_aroussi@hotmail.com  

Benaoumeur Aour 

Laboratory of Applied Biomechanics 
and Biomaterials (LABAB) 
ENPO-MA, Oran, Algeria  

benaoumeur.aour@enp-oran.dz 

Abdelhadi Sabik Bouaziz 

Laboratory of Mechanical 
Manufacturing and Technology 

ENPO-MA, Oran, Algeria 

bouaziz_sabik@hotmail.com 

 

 

Abstract—The electrochemical behavior of stainless steel and 

titanium alloys is affected after prolonged contact with basic or 

acidic solutions, indicating a change in their surface properties. 

The human body often rejects invasive devices that aim to alter 

the biological or chemical composition of blood or other body 

fluids. Stents, fixation plates and screws, spinal implant devices, 

aneurysm clips, intramedullary nails and stems, temporary 

fixation devices and surgical instruments, etc. have been made 

from stainless steel AISI 316L for several years. Although the 
mechanical performance of implants and devices may be 

governed by their bulk properties, their interaction with the 

environment is managed by the characteristics of their superficial 

layer. In the case of biomedical devices, resistance to corrosion 

and biocompatibility has paramount importance. This study 

compares the corrosion behavior of 316L stainless steel and a 
titanium alloy in a Hank solution. The obtained results show that 

the titanium alloy has a higher potential than 316L stainless steel 
and lower corrosion current. 

Keywords-corrosion; 316L stainless steel; titanium; passivation; 

biocompatibility 

I. INTRODUCTION  

Implants are organs and devices in order to replace a 
missing or a damaged organ of the human body, ensuring its 
vital functions optimally. Biomaterials, one of the great 
therapeutic advances in implantation, work under biological 
stress and play an increasingly crucial role in our health and 
quality of life [1]. In order to protect people against the 
possible dangers and inconveniences of implants and devices, it 
is necessary to test sufficiently their elements without putting 
in danger the patient’s life. This is why large-scale studies and 
clinical trials are required [2]. Experiments carried out over the 
years have shown that cells’ response to the surface of the 
biomaterial varies according to their type and state of 
maturation. In vitro and in vivo tests have led several authors to 
conclude that the attachment of osteoblasts is reinforced by the 
roughness of the surface [3, 4]. The first tests, regarding the 
resistance to corrosion in physiological environment, in vitro 
and in vivo, seem conclusive and should continue in the 
coming years [5]. Research was carried out for orthopedic 
applications, such as knee and hip prostheses [6, 7]. This type 
of prosthesis undergoes significant mechanical frictional 

stresses, causing the release of metal particles in large 
quantities. 

Metallic biomaterials have been used in various biomedical 
applications for a century, due to their low cost, excellent 
mechanical properties and inertness [8, 9]. Stainless steel and 
titanium alloys are frequently employed in many biomedical 
applications, including cardiovascular stents/valves, orthopedic 
prosthesis and other devices and implants used in biomedicine, 
because of their biocompatibility and mechanical properties, 
such as high strength and toughness [10, 11]. Titanium alloys 
are extensively used due to their attractive combination of high 
strength, low density, and good corrosion resistance [12,13]. 
Moreover, the 316L stainless steel is known for its good 
ductility, fatigue resistance, and high strength toughness [14, 
15]. However, there is always a concern about their corrosion 
resistance in physiological medium and their biocompatibility 
[16, 17]. The effects of surface treatment and metallic coating 
on the corrosion behavior and biocompatibility of surgical 
316L stainless steel implants were evaluated in [18]. The 
experimental results indicated that Nb coating and surface 
treatment of the stainless steel improved its corrosion behavior. 
The structural and bio-corrosion barrier performance of Mg-
substituted fluorapatite coating on a 316L stainless steel human 
body implant was investigated in [19], concluding in improved 
corrosion resistance and biocompatibility. In [20] the 
microstructural evolution, mechanical, sliding wear and 
corrosion behavior of 316L stainless steel multi axially forged 
at 600°C was studied, showing that severely deformed ultra-
fine-grained metals have the potential to deliver improved 
implant performance. The electrochemical tests conducted in 
[21] showed that the surface enriched by molybdenum 
improved the corrosion resistance of the 316L stainless steel. 

Recently, a comparative assessment of the corrosion 
intensity of orthodontic arch wires made of alloy steel, nickel-
titanium and titanium-molybdenum alloys, in laboratory 
conditions, was presented in [22]. The highest resistance to 
corrosion was observed in nickel-titanium alloy arch wires, 
while steel wires had the lowest. Surface modification allows 
improvements on the materials’ characteristics without 
compromising their important bulk properties. A review on 
stainless steel surface modification methods provides guidance 

Corresponding author: Djelloul Aroussi 



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for selecting the appropriate modification routes, tailored for 
specific biomedical applications [23]. 

The purpose of this study is to compare the corrosion 
behavior of 316L stainless steel and titanium alloy in a Hank's 
solution, utilizing electrochemical measurements, such as 
potentiodynamic and cyclic polarization. 

II. STUDIED MATERIALS 

AISI 316L stainless steel and titanium alloys are commonly 
used in the manufacturing of prostheses. The chemical 
composition of these materials was determined by 
spectroscopic analysis (Tables I and II). Samples were cut in 
the form of a 14mm in diameter disc from cylindrical bars 
delivered in a raw state. 

TABLE I.  COMPOSITION OF 316L STAINLESS STEEL (IN MASS %) 

C P S Si Mn Cr Ni Mo Ti N Cu Co 

0.024 0.029 0.025 0.407 1.519 16.64 10.355 2.037 0.006 0.047 0.296 0.188 
 

TABLE II.  COMPOSITION OF THE TITANIUM ALLOY (IN MASS %). 

TI Fe V Pd 

98.44 0.22 1.33 0.01 
 

III. SURFACE PROPERTIES AND POLISHING TECHNIQUES 

Optimizing the properties of surface corrosion resistance is 
a wide and important research topic, given its potential 
application in different biomedical fields. Corrosion resistance 
is one of surfaces’ major characteristics that must be optimized 
in order to improve the integration of implants. In this section, 
focus is set on roughness and electrochemical properties. 
Anodic dissolution of the sample can produce excellent 
polishing results without deforming its surface. The sample is 
polarized at the anode and connected to the cathode through an 
electrolyte bath concentrated in acid. Applying voltage creates 
a current that travels through the bath and thus preferentially 
alters the surface to be polished. Previously, the surface of the 
sample was mechanically pre-polished on abrasive papers of 
decreasing grain size (up to 1500 grains/cm

2
). It should be 

noted that for each electro-polishing treatment, a new fresh 
electrolyte was used because changes in the ionic metal 
concentration could have an influence on the electro-polishing 
states [24]. Consequently, the sample was cleaned by 
immersion in a mixture of hydrofluoric acid (2% v/v), nitric 
acid (10% v/v) and DI water (88% v/v), for 30s at 50°C, in 
order to dissolve the salts without attacking the metal [25]. 
After each operation, the sample was ultrasonically cleaned in a 
solution of water and acetone for 10 minutes and dried with 
compressed air. The roughness was measured using a “Surftest 
201” portable rugosimeter. Each measurement was repeated 
five times at different locations and the roughness values of the 
samples are recorded in Table III. 

TABLE III.  SURFACE ROUGHNESS OF THE STUDIED MATERIALS 

 Raavg [µm] 

316L (electrolytically polished) 0.05±0.02 

Titanium alloy (mechanically polished) 0.07±0.02 

 

IV. RESULTS AND DISCUSSION 

A. Microhardness  

The microhardness tests were carried out under a load of 
20g, load from which the impression is visible under the 
microscope, using a Vickers microdurometer equipped with a 
microscope. Vickers HV hardness is related to the length of the 
diagonal of the impression (d) and the applied load (m) by: 

2
1.854 /HV m d=  (1) 

with m in [kgf] and d in [mm]. The microhardness results for 
both materials are shown in Table IV and represent the average 
of 10 measurements. It can be noted that the titanium alloy is 
mechanically stronger than 316L stainless steel. 

TABLE IV.  MICROHARDNESS OF THE TESTED SURFACES OF THE 
STUDIED MATERIALS 

Materials Microhardness 

316L stainless steel 166.8 ±0.2Hv 

Titanium alloy 354.0 ±0.2Hv 

 

B. Electrochemical Analysis 

1) Used Solution 

The Hank solution was used as an aggressive environment, 
assimilated to the physiological medium prepared from the 
compounds given in Table V. 

TABLE V.  CHEMICAL COMPOSITION OF THE HANK SOLUTION 

Compositions Quantity (g/l) 

NaCl 8 

CaCl2 0.15 

KCl 0.40 

NaHCO3 0.35 

Glucose 0.1 

MgCl2.6H2O 0.1 

Na2HPO4.2H2O 0.6 

KH2PO4 0.06 

MgSO4.7H2O 0.06 

 

2)  Experimental Conditions 

The samples were characterized electrochemically under 
the following conditions: 

• A simulated physiological environment (Hank's solution of 
PH = 7.1) 

• Environment temperature: 37°C (human body temperature); 

• Sweep rate of the potential: 0.5 mV/s 

3) Evolution of the Free Potential 

This technique makes possible to follow the evolution over 
time of the sample potential (working electrode) in an open 
circuit (OCP) relative to a reference electrode. The sample was 
kept immersed in the electrolyte and the free electrode potential 
was measured as a function of time. The abandonment 
potential (Ea) is determined from the potential follow-up 
curves as a function of time. This potential is a characteristic of 
the modification of the interface between a metal and its 



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environment. This simple technique provides preliminary 
information on the nature of the processes developed on the 
metal/electrolyte interface: corrosion and passivation. The 
general shape of the curve indicates whether the material 
becomes passive if there is an increase in potential as a 
function of time. So, free corrosion tests were performed for all 
treatment conditions [28]. The results shown in Figure 1 
indicate that in a Hank’s solution the native passive films of 
electrolytically polished 316L steel have greater stability. 
Indeed, during 1 hour of immersion the potential of 316L 
polished by electrolysis remains stable at -63mV/ECS, in 
agreement with [26]. 

 

 
Fig. 1.  Evolution of free potential as a function of time of 316L steel in 

Hank's solution 

The curve shown in Figure 2 displays an increase in 
potential over time. In this case, the value of the potential is 
rapidly stabilized, in less than 30 minutes, at about  
-10 mV/ECS. In addition, the anodic and cathodic reactions are 
rapidly balanced. 

 

 
Fig. 2.  Evolution of the free potential as a function of time of the titanium 

alloy in Hank's solution 

The variation of the potential value with immersion time 
can be explained by the electrochemical reaction which begins 

to take place between the medium (Hank's solution) and the 

exposed surface of the metal substrate, more or less adherent, 

causing a relative stability of the potential for corrosion 
abandonment. 

TABLE VI.  FREE POTENTIAL RESULTS FOR STUDIED MATERIALS 

 E(free) (mV/ECS) 

316L stainless steel -63 

Titanium alloy -10 

 

4) Potentiodynamic Curves log(I)=f(E) 

A metal immersed in any electrolytic medium tends to 
dissolve and electrically charge creating a double 
electrochemical layer, comparable to an electric battery. After a 
long enough time, a steady state is established and the metal 
electrode takes, relative to the solution, a potential called the 
potential corrosion (Ecor). The latter cannot be determined 
absolutely and it is marked with respect to a reference 
electrode. Using an external generator and a counter-electrode, 
a current passes through the metal electrode, its steady state is 
modified and its surface takes a new potential value (the curves 
E=f(I) or I=f(E) constitute the polarization curves). Figures 3 
and 4 show the evolution of the polarization curve log(i) in 
terms of the potential E for both materials. The potential 
scanning range was set at a speed of 0.5mV/s. Calculation of 
the density icor and corrosion potential Ecor is performed by 
extrapolating the anode and cathode tafel lines to the corrosion 
potential Ecor. The point of intersection gives us directly icor 
and Ecor. The corrosion rate Vcor, expressed in (mm/year), is 
calculated by [27]: 

( )10 . .

. .

cor

cor

i t M
V

n E d
=  (2) 

where 
cor
i is the corrosion current density (A/cm

2
), t is the time 

corresponding to years, M is the average atomic mass of the 
studied material, n is the valence, and d is the density of 
materials. 

In Figure 3, it is observed that the titanium alloy in the 
solution becomes passive and has a passivation plateau of the 
order of 500mV. The corrosion current is very low and is of the 
order of 0.01µA/cm

2
 (Table VII). So, the titanium alloy has a 

low rate of degradation (60×10
-3
mm/year). In fact, the 

corrosion current can be likened to a residual current resulting 
from an ion diffusion phenomenon through the passive layer 
that has been formed. A break in the passivation layer was 
noticed beyond +500mV/ECS. It should be noted that this 
break was partial because the measured currents remained low. 
Beyond +1.0V/ECS, where the overvoltage is greater, the 
currents increase drastically, meaning that the layer is 
completely broken. This state gives a good behavior against 
corrosion. The 316L steel exhibits the same behavior as the 
titanium alloy. The polarization curve (Figure 4) shows a 
passivation level of +500mV/ECS and a film breaking potential 
of +480mV/ECS. The corrosion current remains low, which is 
of the order of 2µA/cm

2
 and the resistance to polarization is 

weak (10.7MΩ) (Table VIII). 

According to the results obtained in the Hank solution for 
both alloys, significant differences were noted in their 
electrochemical characteristics values. The corrosion potential 
of the titanium alloy is close to -190.8mV/ECS and its 
corrosion current density was 9.8×10

-3
µA/cm

2
. Compared to 

the 316L stainless steel values, titanium’s alloy values are 



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clearly favorable. This alloy has greater potential than 316L 
stainless steel and lower corrosion current. The corrosion rate 
for the titanium alloy is 60×10

-3
mm/year. This value is very 

acceptable in the corrosion field, showing better corrosion 
behavior in the Hank solution (similar to physiological 
environment) compared to 316L stainless steel. Titanium-based 
alloys and stainless steels are passive materials sensitive to 
pitting corrosion in the presence of specific aggressive 
materials. These are most often chloride ions (Cl

-
), which are 

the most aggressive and play an important role in the pitting 
mechanism.  

 
Fig. 3.  Potentiodynamic curve of the titanium alloy in a Hank solution 

TABLE VII.  POTENTIODYNAMIC RESULTS FOR TITANIUM ALLOY 

Corrosion 

parameter 

Ecorr  

(mV/ECS) 

Icorr 

(µA/cm
2
) 

Vcorr 

(mm/yr) 

Rp 

(MΩ) 

Titanium alloy -190.8 0.0098 60×10
-3
 2.21 

 

 
Fig. 4.  Potentiodynamic curve of 316L stainless steel in a Hank solution 

TABLE VIII.  POTENTIODYNAMIC RESULTS FOR 316L STEEL 

Corrosion 

parameter 

Ecor  

(mV/ECS) 

Icor 

(µA/cm
2
) 

Vcor 

(mm/yr) 

Rp 

(MΩ) 

316L stainless steel -267.5 2.155 16.86×10
-3
 10.7 

 

Pitting corrosion is related to localized attacks of 
heterogeneities that contribute to the deterioration of the 
passive layer protecting the material. This process of material 
deterioration generally leads to a generalized corrosion 
phenomenon. Observation by an optical microscope of the 

samples for both studied alloys shows the attacking shape they 
experienced after a corrosion test in Hank's solution (Figure 5). 

 

(a) 

 

(b) 

 

Fig. 5.  Illustration of pitting corrosion in the case of (a) titanium alloy and 

(b) 316L stainless steel 

5) Cyclic Potentiodynamic Curves 

The behavior of the oxide layer formed on the samples’ 
surface was studied by drawing the cyclic potentiodynamic 
curves. Figure 6 shows the titanium’s alloy cyclic polarization 
curves over a potential range for a half-hour immersion time in 
Hank's solution with a sweep rate of 0.5mV/s. There is a 
passivation stage with a slight increase in the corrosion 
potential compared to the previous value of EC for the “go” 
scan. On the other hand, during the “return” scan the corrosion 
potential is less great with higher current densities. This means 
that the formed oxide layer has instabilities. This degradation 
may be at the source of localized corrosion.  

 

 
Fig. 6.  Cyclic potentiodynamic polarization curve of titanium alloy 

Cyclic curves in the case of 316L (Figure 7) indicate a 
different behavior. Beyond the formation of the passivation 
layer on this alloy, an ennoblement of the potential corrosion in 



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the anodic direction of the order of 300mV and very low 
corrosion currents were noted. Stable protective character can 
be established here. Tables IX and X summarize the cyclic 
potentiodynamic results for titanium alloy and 316L stainless 
steel. It is noted that the polarization resistance of the titanium 
alloy is higher than that of 316L. 

TABLE IX.  CYCLIC POTENTIODYNAMIC RESULTS FOR THE 
TITANIUM ALLOY 

Corrosion 

parameter for the 

Titanium alloy 

Ecor 

(mV/ECS) 

Icor 

(µA/cm
2
) 

Vcor 

(mm/year) 

Rp 

(KΩ) 

-419.0 0.347 53.62×10
-3
 62.51 

 

 
Fig. 7.  Cyclic potentiodynamic polarization curve of 316L stainless steel. 

TABLE X.  CYCLIC POTENTIODYNAMIC RESULTS FOR THE 316L 
STEEL 

Corrosion 

parameter for 

316L stainless steel 

Ecor 

(mV/ECS) 

Icor 

(µA/cm
2
) 

Vcor 

(mm/year) 

Rp 

(KΩ) 

-38.5 3.92 30.65×10
-3
 5.54 

 

V. CONCLUSION 

This work compared the behavior of two passive metals in 
a biological medium, the 316L stainless steel and a titanium 
alloy. Several experimental techniques were used, starting with 
microstructural characterization and microhardness 
measurements for both materials. Electrochemical tests allowed 
the study of the biological medium’s effects on the 
electrochemical behavior of the materials. The systems were 
studied by polarization curves, and were distinguished by 
different responses, proving that the biological medium has a 
great influence on the electrochemical behavior of both 
materials. Due to the importance of biomaterials applications, 
especially in the biomedical field, the samples were immersed 
in a Hank’s solution and the evolution of the surface state was 
examined over time. The results obtained can be summarized 
as: 

• Microhardness measurements show that the titanium alloy 
is mechanically stronger than 316L stainless steel 

• The steady-state potentiodynamic curves determine 
equilibrium potential for both alloys. 

• In the Hank’s solution significant differences were found in 
the electrochemical characteristics. The titanium alloy had 
higher potential than 316L stainless steel and lower 
corrosion current 

• The use of cyclic intensity-potential curves, can help 
determining the quality of the passive film 

• Finally, the study of the polarization resistance evolution as 
a function of the immersion time shows that the two alloys 
retain their resistance properties as a function of the 
immersion time in the Hank’s solution. 

ACKNOWLEDGMENT 

The authors thank the laboratory of the ALFON complex 
(Algerian Foundry Society of Oran) for their support and 
assistance in the determination of chemical compositions by 
spectroscopic analysis. 

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